The ability to detect various physical and electrical activities along with the presence of certain substances in certain systems is desirable for not only medical diagnosis but also for process control, environmental monitoring, and for general real-time chemical and biological analysis. Current optical sensing approaches for biological and chemical measurements are based on fluorescence and surface plasmon resonance (SPR) techniques. However, these techniques have been found to be cumbersome and lack in sensitivity. In addition, these approaches fail to account for measuring certain physical and electrical changes which may be occurring in the system.
Optical sensors based on fluorescence require multiple fibers or tapering of an optical fiber prior to applying an affinity coating. Light propagating in the tapered region of the optical fiber has an evanescent field that extends into the coating. A sample having an unknown concentration of specific biological agents is doped with a fluorescent dye that is excited by the evanescent signal wavelength. The specimen is adsorbed by the coating and the fluorescent signal of the biological agents is coupled back into the fiber. The magnitude of the fluorescent signal is an indication of the specimen concentration.
It has been found that the manufacturing of these sensors is labor-intensive and not cost-effective because the optical fiber must be tapered. There is no alternative to tapering the fiber because the wavelength of the excitation and fluorescent signal are such that propagation of the fluorescent signal in the fiber is lossy. Tapering the optical fiber reduces some of this loss mechanism but weakens the structural integrity and thus, severely limits field use. In addition, these sensors have limited signal-to-noise ratios due to the background fluorescence of non-bound agents. Since the interrogation of the device determines fluorescent intensity, fluctuations due to background signals create errors in the measurement. In turn, the sensitivity of these sensors is much less than the sensitivity of the sensors of the present invention.
Bulk SPR-based techniques are difficult to transition into a mobile platform because critical component alignment cannot be maintained in a field environment. To ruggedize the system, impractical size and weight restrictions are required. Optical fiber-based SPR sensors have been made but they have many limitations. In SPR, an optical transverse magnetic (TM) wave is coupled to a surface wave which is created in a metallic layer such as gold. SPR wave excitation occurs due to the evanescent field created by an incident wave totally internally reflected (TIR) at a metal interface. The energy level of the metal electron gap spacing must correspond to the excitation wavelength of the evanescent field. If TIR or other stimulation properly matches the momentum and propagation direction of the incident and SPR signal, excitation will occur and hence absorption of that particular wavelength. When it is desirable to determine biological or chemical concentration, an affinity coating is applied to the non-excitation side of the metallic layer. When adsorption occurs, the refractive index of the sample changes which results in a change in the surface plasmon excitation conditions. The shift in the absorption wavelength is then monitored to determine refractive index changes which are correlated to chemical or biological agent concentration.
Single mode fiber, when used for SPR, does not guarantee transverse magnetic (TM) light propagation. Thus the polarization of light in low-birefringence fiber needs to be adjusted until an SPR wave is excited. Although such polarization control works well in the laboratory environment, in field applications the polarization is difficult to maintain when the fiber is perturbed during sensing. When multimode fiber is used, both TM and transverse electric (TE) modes are propagated. However, due to the geometry of the propagating modes, the maximum coupling of TM light is 50% or 3 dB. Thus, only one-half of the input light can possibly excite an SPR wave. The isolation of the absorption band of the SPR signal is a theoretical maximum 3 dB, and in practical systems it is typically 1.8 dB due to the relatively small evanescent fields in lower-order modes. The isolation of the absorption band of a long period grating, used in the present invention, is typically 25 dB down from the peak spectral signal. When specifically designed optical fiber and gratings are used, the isolation of the absorption band has a potential of greater than 35 dB down from the peak spectral signal.
A mode scrambler is required for SPR sensors to ensure that all modes are excited equally. Each mode will have a different sensitivity and when a spectral shift occurs due to the refractive index change of a test specimen, the spectral shape of the absorption signal changes. This shape change leads to detection errors through misinterpretation of absorption band location. Optical waveguide sensors based on long period gratings generate significantly larger signal-to-noise ratios allowing for the determination of the long period grating coupling wavelength to be more accurate and accomplished with simpler demodulation electronics. Moreover, the shape of the long period grating does not change when shifting.
SPR sensors, in general, operate at the same wavelength for similar biosample evaluation. Complex optical or mechanical switching is needed to monitor multiple sensors with a single source/detector system. This reduces the reliability of the device by introducing additional mechanical elements as well as increasing the overall cost of the system. The long period grating sensors of the present invention couple at wavelengths determined by the long period gratings written in the waveguide using high energy ultraviolet light. These periodic index variations are written with different spacings to long period gratings of different wavelengths. The long period grating can therefore be demodulated using standard wavelength division multiplexing techniques. However, SPR sensors and fluoroscopy sensors are limited to specific absorption and fluorescent wavelengths.
The long period grating based sensors of the present invention are easily manufactured. When SPR optical fiber probes are produced, the cladding must be stripped to precise dimensions. The region must be radially symmetric throughout so the light within the fiber maintains a constant propagation angle. Next, a metallic layer must be coated onto the stripped region at a thickness that is radially symmetric. If the thickness of the metallic layer varies, the SPR coupling conditions change and thus reduce sensitivity. The precision involved in making these sensors makes it difficult to manufacture them in large quantities. The long period grating based sensors of the present invention are made consistently and accurately through exposure to a high energy UV signal. The signal is generated by illumination of an amplitude mask.
Tran et al. ("Real-Time Immunoassays Using Fiber-Optic Long-Period Grating Sensors", Biomedical Sensing, Imaging, and Tracking Technologies I, Proceedings SPIE--The International Society for Optical Engineering, R. A. Lieberman et al., Eds., Vol. 2676, Jan. 29-31, 1996, pp. 165-170) demonstrated how a long-period grating was used to detect and monitor in real time the interaction of a specific antigen to an immobilized antibody on a silica fiber. Long-period gratings optically couple the fundamental guided mode to discrete cladding modes that attenuate rapidly on propagation along the length of an optical fiber. Any change in the properties of the cladding or the material surrounding the cladding, modifies the coupling mechanism and results in a modulation of the output optical spectrum. If a broadband light source is injected into the optical fiber, the attenuation spectra arising from the coupling to different cladding modes suitably modifies the original spectral profile. Long period grating based biosensors utilize the sensitivity to the index of refraction variation in the medium enclosing the cladding as a means of monitoring serological interactions. In their experiments, the antibody was adsorbed to the silica optical fiber. As toxin A complexes with the antibodies, the molecular length of the entire compound increases and the density of the bound molecules increases thus changing the refractive index.
However, this system was found to be deficient and the test results are questionable. Unfortunately, Tran et al. failed to run a control so there is no way to determine what caused the change in wavelength. In turn, the changes in output may have resulted from non-bonding interactions. When the antibodies are adhered directly to an optical fiber, the antibody density is limited by the antibody size. The density affects the number of target molecules that can be captured and also the refractive index of the coating. To obtain measurable sensitivity from the antibodies, the population of attached antibodies must be as large as possible. This is counter-productive because the large population results in the refractive index of the coating becoming too large to be useful. Antibodies have two sites for binding to target molecules. When antibodies are applied to an optical fiber, they are packed too close together causing the binding sites of adjacent antibodies to be positioned close together. This results in the overlap of binding sites, which precludes potential binding with the target molecules. The close packing of the binding sites prevents capture by either site. In addition, many of the binding sites are oriented toward the fiber instead of toward the sample, reducing the overall number of binding sites. Lastly, the target molecule (protein) that the antibodies on the fiber are trying to capture is typically greater than the antibody spacing. In turn, antibodies end-up competing for capture and no binding occurs with the coating.
An object of the present invention is to provide an optical sensor having a reactive coating.
Another object of the present invention is to provide an optical sensor having a waveguide and a long period grating wherein a reactive coating which is positioned in an operable relationship to the long period grating such that the reactive coating causes the long period grating to produce a wavelength transmission spectrum functionally dependent on the parameter sensed.
Another object of the invention is to provide an optical sensor which is more sensitive than SPR or fluoroscopy.
Another object of the invention is to provide an optical sensor which does not require metal on a waveguide surface.
Another object of the invention is to provide an optical sensor which has at least one long period grating disposed within an optical waveguide and a reactive coating which is either physically reactive, electrically reactive or chemically reactive.